It is standard practice to fix implants, in particular endo-joint implants into bone by means of an autopolymerising two-component bone cement prepared by mixing liquid methylmethacrylate (MMA) with polymethylmethacrylate (PMMA) in powder form.
The use of polymethylmethacrylate (PMMA) cement in total hip replacement has been extremely successful since its introduction by Charnley in 1960. However, many revisions are still required every year and a large majority of these are due to aseptic loosening (Stauffer et al 1982; Sutherland et al 1982; Fournasier et al 1976). In particular, it is the loosening of the femoral component which accounts for the majority of such cases (Poss et al 1970; Jasty et al 1991). Failure of the cement mantle is believed to initiate at the stem-cement interface (James et al 1993; Culleton et al 1993; Jasty et al 1991) which is significantly weaker than the bulk cement under static and fatigue loading (Raab et al 1981). Furthermore, a number of studies indicate that the stresses become significantly higher in the cement mantle once the stem-cement interface has become loose (Harrigan et Harris 1991; Crowninshield et Tolbert 1983; Freitag et Cannon 1977).
Both the static and fatigue strength of bone cement decrease with porosity (James et al 1992; Burke et al 1984; Saha et al 1984; DeWijn et al 1975; Greenwald et al 1977) and the stem-cement interface strength is similarly affected (Welsh et al 1971, James et al. 1993). Fatigue is the most credible mode of mechanical failure of the cement mantle (Harris et Davies 1988) and any measure to reduce the porosity of both the bulk cement and its interface to the prosthesis should be of clinical benefit.
Pores in the bulk cement result primarily from air bubbles which become entrapped during hand mixing of the powder and liquid components. It is also thought that monomer evaporation at the high temperatures of polymerisation due to its volatility may contribute to pore formation (Debrunner et al 1976). The bubbles act as stress risers of at least factor two (Timoshenko & Goodier 1934; Burke et al 1984; Freitag et al 1977; Carter et al 1982; Gates et al 1983) and it has been reported that failure in fatigue has occurred in test specimens almost exclusively through a pore (cited in James et al 1992). Reduction in porosity has been achieved using two techniques: centrifugation (Burke et al 1984) and mixing under vacuum (Schreurs et al 1988; Lidgren et al 1987). Other methods to decrease the porosity include pressurising cement (Saha et Pal 1984) and the use of improved hand mixing technique (Eyerer et Jin 1986).
The stem-cement interface strength can be improved by precoating of the stem with a thin layer of PMMA under optimal polymerisation conditions (Ahmed et al 1984). Under static testing of the precoated interfaces the bulk PMMA fractured preferentially to the stem-cement interface, thus implying an interface strength in excess of the bulk cement strength. Roughening the implant surface, both on the micro and macro scale, improves the interface strength, but it also increases stresses in the cement mantle (Stone et al 1989; Welsh et al 1971).
The bone-cement interface appears to be less critical in loosening and is believed to be due to biological factors rather than mechanical (Schmalzried et al 1992). There are a number of practised techniques which are used to ensure an optimal cement-bone interface (Krause et al 1982). These include thorough cleaning of the canal using pressurised water picks, and the pressurisation of cement into the cancellous bone interstices (Bourne et al 1984). This ensures maximum area for load transfer. However, Noble and Swarts (1983) have shown that a penetration of more than 5 mm is likely to cause thermal damage to the bone and from mechanical considerations more than 2.9 mm is possibly superfluous (Jansson 1993).
The heat of polymerisation of PMMA cement is considered to be an important factor in initial bone necrosis (Dipisa et al 1976). Lenhartz (1959) cites the threshold temperature of collagen denaturation (56.degree. C.) as the criterion for bone necrosis. However, Moritz and Henriques (1947) present a time dependant threshold level for cell necrosis (e.g. a temperature of 50.degree. C. for three minutes will kill bone cells). Toksvig-Larsen et al. (1991) measured a mean maximum temperature of 40.degree. C. at the cement-bone interface using lavage and vacuum mixed cement. They found that cooling of the stem prior to implantation had no affect on the peak cement temperature and indeed this increases the setting time of the cement(De Waal Malefijt et al 1987) and compromises its mechanical properties.
Porosity at the stem-cement interface is caused by shrinkage of the cement. If the stem is implanted at room temperature the cement at the warmer bone-cement interface will polymerise first, creating a stiff shell towards which cement will shrink. Thus, the cement shrinks away from the stem, creating pores at the areas of least adhesion.